X-ray imaging is a powerful tool in many fundamental and practical applications. As a primary example, X-ray computed tomography (CT) is a cornerstone of modern hospitals and clinics. The dominating theory of X-ray imaging is generally based on the attenuation contrast mechanism.
X-ray gratings have been used for hybrid CT imaging in terms of attenuation, refraction, and small-angle scattering. This grating-based approach represents a paradigm shift in X-ray CT from gray-scale (attenuation) to true-color (attenuation, refraction, and small-angle scattering, which is also referred to as dark-field, and spectral) imaging.
In conventional X-ray imaging, the image contrast arises from varying linear attenuation coefficients. Attenuation-contrast-based imaging exhibits good performance only when strong attenuators are embedded in a weakly absorbing matrix, such as in the cases of bone-tissue and tissue-air interfaces. However, biological soft tissues include mainly light elements (e.g., hydrogen, carbon, nitrogen, and oxygen), and their compositions are quite homogeneous with little density variation. The attenuation-contrast between soft tissue features is often insufficient to reflect pathological changes.
In particular, many healthy tissues display similar characteristics in current X-ray images as those of diseased tissues, such as tumors. For example, fibro-glandular tissue can have a density of 1.035 cm−3 and an attenuation coefficient of 0.80 cm−1, and cancerous tissue can have a density of 1.045 cm−3 and an attenuation coefficient of 0.85 cm−1. Given inherent measurement noise, it is challenging to discern such cancerous tissue from the healthy tissue, as well as other soft tissue features such as those reflecting musculoskeletal healing. Therefore, attenuation-contrast-based imaging is unable to differentiate early-stage tumors and soft tissues.
X-ray small-angle scattering, or dark-field, imaging acquires photons slightly deflected from the primary beam through an object. Small-angle scattering signals reflect structural texture on length scales between 1 nanometer (nm) to several hundred nm. This imaging mode can reveal subtle texture of tissues. For example, the growth of tumors causes remarkable differences in small-angle scattering patterns from that of healthy tissues. It is clinically important that the structural variation in a tumor modifies the refractive index. The propagation of X-rays in a medium is characterized by the complex index of refraction. The cross section of an X-ray phase shift is one thousand times larger than that of the linear attenuation in the 20 keV-100 keV range. This implies that phase-contrast imaging has much higher sensitivity for light elements than does attenuation-contrast imaging.
The contrast-to-noise ratio of differential phase contrast CT images is superior to that of its attenuation-contrast CT counterpart. Therefore, phase-contrast imaging can observe unique critical structures of soft biological tissues. Moreover, the refractive index of tissues is inversely proportional to the square of the X-ray energy while the absorption coefficient decreases as the fourth power of the X-ray energy. Hence, X-ray phase-contrast imaging is suitable to operate at higher energies (e.g., >30 keV) for lower radiation dose than is attenuation imaging. Higher energy X-ray imaging can be useful for imaging and studies on large animals and/or human patients.
In X-ray grating-based imaging, X-ray phase-shift and dark-field information is currently extracted using the fringe scanning method through shifting of the analyzer grating. However, major challenges include the difficulty in making an analyzer grating (often known as G2), the requirement of a high-precision mechanical device, and the long data acquisition time due to the use of the phase stepping procedure.